Classifications

• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
  • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
  • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
  • G01R33/385—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using gradient magnetic field coils
  • G01R33/3856—Means for cooling the gradient coils or thermal shielding of the gradient coils
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
  • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
  • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
  • G01R33/3802—Manufacture or installation of magnet assemblies; Additional hardware for transportation or installation of the magnet assembly or for providing mechanical support to components of the magnet assembly
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
  • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
  • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
  • G01R33/381—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using electromagnets
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
  • G01R33/44—Arrangements or instruments for measuring magnetic variables involving magnetic resonance using nuclear magnetic resonance [NMR]
  • G01R33/445—MR involving a non-standard magnetic field B0, e.g. of low magnitude as in the earth's magnetic field or in nanoTesla spectroscopy, comprising a polarizing magnetic field for pre-polarisation, B0 with a temporal variation of its magnitude or direction such as field cycling of B0 or rotation of the direction of B0, or spatially inhomogeneous B0 like in fringe-field MR or in stray-field imaging
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
  • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
  • G01R33/38—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field
  • G01R33/385—Systems for generation, homogenisation or stabilisation of the main or gradient magnetic field using gradient magnetic field coils
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
  • G01R33/00—Arrangements or instruments for measuring magnetic variables
  • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
  • G01R33/28—Details of apparatus provided for in groups G01R33/44 - G01R33/64
  • G01R33/42—Screening
  • G01R33/421—Screening of main or gradient magnetic field

Definitions

• the present invention relates generally to magnetic resonance imaging. More specifically, the present invention relates to an arrangement of coils for increasing signal detection sensitivity of a magnetic resonance imaging system.
• Magnetic resonance imaging is a major imaging technique used in medicine. MRI is capable of generating detailed images of soft tissues such as the brain, muscles and kidneys. Specific properties of the various compounds found inside tissues, such as water and/or fat, are used to generate images.
• MRI Magnetic resonance imaging
• the vector sum of the nuclear magnetic moments of a large number of atoms possessing a nuclear spin angular momentum, such as hydrogen, which is abundant in water and fat will produce a net magnetic moment in alignment with the externally applied field.
• the resultant net magnetic moment can furthermore precess with a well-defined frequency that is proportional to the applied magnetic field. After excitation by radio frequency pulses, the net magnetization will generate a signal that can be detected.
• Delta relaxation enhanced magnetic resonance generally referred to as field-cycled relaxometry or field-cycled imaging is an MRI technique that offers the possibility of using underlying tissue contrast mechanism which vary with the strength of the applied magnetic field to generate novel image contrasts.
• DREMR contrast the main magnetic field is varied as a function of time during specific portions of an MR pulse sequence.
• a field-shifting electromagnet coil is used to perform the field variation. Proper arrangement of the field-shifting electromagnet with the traditional MRI coils used in a DREMR system is important since the contrast mechanism for DREMR is highly correlated with the strength of the magnetic field shifts produced.
• an integrated magnet device for use in a magnetic resonance imaging (MRI) system.
• the integrated magnet device can comprising field-shifting electromagnets including primary field-shifting magnets and field-shifting shield magnets, the primary field shifting magnets placed closer to an imaging volume than the field-shifting shield magnets.
• the integrated magnet device can further comprise gradient coils that can be placed between the primary field-shifting magnets and field-shifting shield magnets and at least one substrate layer that can provide mechanical support for the field-shifting electromagnets and the gradient coils.
• the integrated magnet device can also include at least one cooling mechanism.
• FIG. 1 shows a block diagram of functional subsystems of a delta relaxation magnetic resonance imaging system in accordance with an implementation
• FIG. 2 shows an imaging volume and corresponding slice to be scanned by the delta relaxation magnetic resonance system of FIG. 1 in accordance with an
• FIG. 3 shows illustrative examples of T1 and T2 relaxation diagrams
• FIG. 4 shows an example pulse sequence in accordance with an
• FIG. 5 shows a schematic representation of a k-space containing one received line in accordance with an implementation
• FIG. 6 shows an idealized radial cross-section of an example integrated magnet device in accordance with an implementation
• FIG. 7 shows idealized longitudinal cross-section of an example integrated magnet device in accordance with an implementation.
• FIG. 1 a block diagram of a delta relaxation magnetic resonance imaging (DREMR) system, in accordance with an example implementation, is shown at 100.
• DREMR delta relaxation magnetic resonance imaging
• Traditional magnetic resonance imaging (MRI) systems represent an imaging modality which is primarily used to construct pictures of magnetic resonance (MR) signals from protons such as hydrogen atoms in an object.
• MR magnetic resonance
• typical signals of interest are MR signals from water and fat, the major hydrogen containing components of tissues.
• DREMR systems use field-shifting magnetic resonance methods in conjunction with traditional MRI techniques to obtain images with different contrast than is possible with traditional MRI, including molecularly-specific contrast.
• the illustrative DREMR system 100 comprises a data processing system 105.
• the data processing system 105 can generally include one or more output devices such as a display, one or more input devices such as a keyboard and a mouse as well as one or more processors connected to a memory having volatile and persistent components.
• the data processing system 105 can further comprise one or more interfaces adapted for communication and data exchange with the hardware components of MRI system 100 used for performing a scan.
• example the DREMR system 100 can also include a main field magnet 1 10.
• the main field magnet 1 10 can be implemented as a
• the main field magnet 1 10 is operable to produce a substantially uniform main magnetic field having a strength B0 and a direction along an axis.
• the main magnetic field is used to create an imaging volume within which desired atomic nuclei, such as the protons in Hydrogen within water and fat, of an object are magnetically aligned in preparation for a scan.
• a main field control unit 1 15 in communication with data processing system 105 can be used for controlling the operation of the main field magnet 1 10.
• the DREMR system 100 can further include gradient magnets, for example gradient coils 120 used for encoding spatial information in the main magnetic field along, for example, three perpendicular gradient axis.
• the size and configuration of the gradient coils 120 can be such that they produce a controlled and uniform linear gradient.
• three paired orthogonal current-carrying primary coils located within the main field magnet 110 can be designed to produce desired linear-gradient magnetic fields.
• the gradient coils 120 may be shielded and include an outer layer of shield magnets, for example coils which can produce a counter magnetic field to counter the gradient magnetic field produced by the primary gradient coils forming a primary-shield coils pair.
• the "primary" coils can be responsible for creating the gradient field and the "shield” coils can be responsible for reducing the stray field of the primary coil outside a certain volume such as an imaging volume.
• the primary-shield coils pair of the gradient coils 120, the primary and shield coils may be connected in series. It is also possible to have more than two layers of coils for any given gradient axis that together form a shielded gradient coil.
• Shielded gradient coils 120 may reduce eddy currents and other interference which can cause artefacts in the scanned images. Since eddy currents mainly flow in conducting components of the DREMR system 100 that are caused by magnetic fields outside of the imaging volume (fringe fields), reducing the fringe fields produced by the gradient coils 120 may reduce interference. Accordingly, the shapes and sizes, conductor wire patterns and sizes, and current amplitudes and patterns of the primary-shield coils pair can be selected so that the net magnetic field outside the gradient coils 120 is as close to zero as possible. For cylindrical magnets, for example, the two coils can be arranged in the form of concentric cylinders whereas for vertical field magnets, the two coils may be arranged in coaxial disks.
• One side effect of shielding can be that the fields produced by the primary- shield coils pair of the gradient coils 120 may partially cancel each other within the imaging volume. Accordingly, more current can be required to produce a gradient field with a particular strength by shielded gradient coils 120 than by unshielded gradient coils 120. This effect can be quantified as the gradient efficiency, which may be defined as the achievable gradient strength for 1 Ampere of driving current.
• Another important parameter describing gradient coil performance is called the gradient slew rate, which is the rate of driving a gradient coil from zero to its maximum amplitude. This term is inversely proportional to the inductance of the gradient coil.
• the inductance in order to increase the efficiency of a shielded gradient coils 120 to be comparable to the efficiency of an unshielded gradient coils 120 the inductance must increase. This increase in inductance will decrease the maximum achievable slew rate.
• the loss in efficiency for a shielded configuration can depend on the distance and current density ratio between the primary and shield coils. Increasing the distance between the primary-shield coils pair may increase the efficiency.
• the conductive components of the gradient coils 120 may consist of an electrical conductor (for example copper, aluminum, etc.).
• the internal electrical connections can be such that when a voltage difference is applied to the terminals of the gradient coils 120, electric current can flow in the desired path.
• the conductive components for the three gradient axes for both the primary gradient coils and the gradient shield coils can be insulated by physical separation and/or a non-conductive barrier.
• the primary gradient windings can be placed on a non-conductive substrate (for example, G10, FR4, epoxy or others).
• the gradient coils 120 may also be provided with thermal control or heat extraction mechanisms.
• some of the windings can be hollow and coolant can be passed through these hollow conductors to extract heat from the gradient coils 120, produced, for instance, by resistive heating of the windings when electricity is applied.
• other methods of extracting heat can be used, such as inserting coolant channels within the gradient coils 120.
• the coolant channels can be in thermal contact with the gradient coil windings.
• the gradient coils 120 can also be mounted in a thermally-conductive but electrically-non-conductive epoxy to ensure that the mechanical assembly is rigid and to limit the possibility of electrical breakdown.
• the magnetic fields produced by the gradient coils 120 can be superimposed on the main magnetic field such that selective spatial excitation of objects within the imaging volume can occur.
• the gradient coils 120 can attach spatially specific frequency and phase information to the atomic nuclei placed within the imaging volume, allowing the resultant MR signal to be reconstructed into a useful image.
• a gradient coil control unit 125 in communication with the data processing system 105 can be used to control the operation of the gradient coils 120.
• the DREMR system 100 there may be additional electromagnet coils present, such as shim coils (traditionally, but not limited to, producing magnetic field profiles of 2nd order or higher spherical harmonics) or a uniform field offset coil or any other corrective electromagnet.
• shim coils traditionally, but not limited to, producing magnetic field profiles of 2nd order or higher spherical harmonics
• a uniform field offset coil or any other corrective electromagnet.
• the corrective electromagnets such as the shim coils, carry a current that is used to provide magnetic fields that act to make the main magnetic field more uniform.
• the fields produced by these coils can aid in the correction of inhomogeneities in the main magnetic field due to imperfections in the main magnet 10, or to the presence of external ferromagnetic objects, or due to susceptibility differences of materials within the imaging region, or any other static or time-varying phenomena.
• the DREMR system 100 can further comprise radio frequency (RF) coils 130.
• the RF coils 130 are used to establish an RF magnetic field with a strength B1 to excite the atomic nuclei or "spins".
• the RF coils 130 can also detect signals emitted from the "relaxing" spins within the object being imaged. Accordingly, the RF coils 130 can be in the form of separate transmit and receive coils or a combined transmit and receive coil with a switching mechanism for switching between transmit and receive modes.
• the RF coils 30 can be implemented as surface coils, which are typically receive only coils and/or volume coils which can be receive and transmit coils.
• the RF coils 130 can be integrated in the main field magnet 1 10 bore. Alternatively, the RF coils 130 can be implemented in closer proximity to the object to be scanned, such as a head, and can take a shape that approximates the shape of the object, such as a close- fitting helmet.
• An RF coil control unit 135 in communication with the data processing system 100 can be used to control the operation of the RF coils 130.
• DREMR system 100 can use field-shifting electromagnets 140 while generating and obtaining MR signals.
• the field-shifting electromagnets 140 can modulate the strength of the main magnetic field. Accordingly, the field-shifting electromagnets 140 can act as auxiliary to the main field magnet 1 10 by producing a field-shifting magnetic field that augments or perturbs the main magnetic field.
• a field-shifting electromagnet control unit 145 in communication with the data processing system 100 can be used to control the operation of the field-shifting electromagnets 140.
• the field-shifting electromagnets 140 may include a shield similar to the shielded gradient coils 120 described above.
• the shielded field-shifting electromagnets 140 can have two components: an inner primary field- shifting electromagnets, to produce the field shift and an outer shield field-shifting electromagnets, to form a shield by reducing the stray field of the primary field-shifting electromagnets outside a certain volume such as an imaging volume.
• one side effect of shielding the field-shifting electromagnets 140 can be that the fields produced by the primary and shield components of the shielded field- shifting electromagnets 140 may partially cancel each other within the imaging volume. Accordingly, more current can be required to produce a magnetic field with a particular strength by shielded field-shifting electromagnets 140 than by unshielded field-shifting electromagnets 140.
• This effect can be quantified as the field-shift efficiency, which may be defined as the field-shift amplitude per 1 Ampere of current passing through the electromagnet.
• the loss in efficiency for a shielded configuration depends on the distance and current density ratio between the shield electromagnets and the primary electromagnets. Increasing the distance between the primary and shield
• the electromagnets may increase the field-shift efficiency.
• the conductive components of the field-shifting electromagnets 140 including the primary and shield electromagnets, may consist of an electrical conductor (for example copper, aluminum, etc.).
• the internal electrical connections can be such that when a voltage difference is applied to the terminals of the field-shifting electromagnets 140, electric current can flow in the desired path.
• the conductive components for both the primary and the shield electromagnets can be insulated by physical separation and/or a non-conductive barrier.
• the field-shift windings can be placed in layers on or within a non-conductive substrate (for example, G10, FR4, epoxy or others).
• the field-shifting electromagnets 140 may also be provided with thermal control or heat extraction mechanisms.
• the windings can be hollow and coolant can be passed through these hollow conductors to extract heat deposited in the electromagnet due to resistive heating of the windings when electricity is applied.
• other methods of extracting heat can be used, such as inserting coolant channels within the field-shifting electromagnets 140.
• the coolant channels can be in thermal contact with the field-shifting electromagnets 140.
• the field-shifting electromagnets 140 can also be mounted in a thermally-conductive but electrically-non-conductive epoxy to ensure that the mechanical assembly is rigid and to limit the possibility of electrical breakdown.
• the DREMR system 100 detects the presence of atomic nuclei containing spin angular momentum in an object, such as those of Hydrogen protons in water or fat found in tissues, by subjecting the object to a relatively large magnetic field.
• the main magnetic field has a strength of BO and the atomic nuclei containing spin angular momentum may be Hydrogen protons or simply protons.
• the main magnetic field partially polarizes the Hydrogen protons in the object placed in the imaging volume of the main magnet 110.
• the protons are then excited with appropriately tuned RF radiation, forming an RF magnetic field with a strength of B1 , for example.
• weak RF radiation signal from the excited protons is detected as an MR signal, as the protons "relax" from the magnetic interaction.
• the frequency of the detected MR signal is proportional to the magnetic field to which they are subjected.
• Cross-sections of the object from which to obtain signals can be selected by producing a magnetic field gradient across the object so that magnetic field values of the main magnetic field can be varied along various locations in the object.
• the variations allow assigning a particular signal frequency and phase to a location in the object. Accordingly, sufficient information can be found in the obtained MR signals to construct a map of the object in terms of proton presence, which is the basis of a traditional MRI image. For example, since proton density varies with the type of tissue, tissue variations can be mapped as image contrast variations after the obtained signals are processed.
• FIG. 2 to further illustrate the example signal acquisition process by the DREMR system 100, it will be assumed that an object is placed within an imaging volume 250 of the main magnet 110 having a main magnetic field 210 with a strength B0, pointing along the Z-axis indicated at 240.
• the object subsequently has a net magnetization vector.
• a slice in a plane along the X and Y axes, as indicated at 205 is being imaged. It should be noted that in this example, the slice has a finite thickness along the Z-axis, creating a volumetric slice 205.
• the individual spins align themselves in the direction of the Z-axis.
• the magnetization by main field B0 can produce a net longitudinal magnetization Mz, with an amplitude of M0, parallel with the main magnetic field.
• Excitation of the spins may be achieved when a radio frequency (RF) pulse that generates the RF magnetic field with an amplitude of B1 is applied at the Larmor frequency, by the RF coils 130.
• RF radio frequency
• B1 applied RF
• magnetization that is projected in the X-Y plane is the net transverse magnetization Mxy.
• the spins can precess about the applied RF magnetic field until the RF magnetic field is removed.
• Transverse relaxation can cause irreversible dephasing of the transverse magnetization. There is also a reversible dephasing effect caused by magnetic field inhomogeneities. These additional dephasing fields may arise from a variety of sources including the main magnetic field inhomogeneity, the differences in magnetic susceptibility among various tissues or materials, chemical shift, and gradients applied for spatial encoding.
• the contribution to the transverse relaxation time from these reversible dephasing effects are typically referred to as T2'.
• the characteristic relaxation time of the combination of reversible ( ⁇ 2') and irreversible (T2) dephasing effects is typically referred to as T2 * relaxation.
• T1 and T2 are important for development of contrast in MR imaging.
• the relaxation times can vary with the strength of the magnetic field applied, as well as temperature.
• T1 and T2 values associated with biological tissues can vary. Generally, tissues with shorter T1 times, such as T1 a as indicated at 315, can yield greater signal intensity at a given point in time (appearing brighter in images) than those with longer T1 times, such as T1 b as indicated at 305, due to the more rapid recovery of signal.
• tissues possessing short T2 times, such as T2a as indicated at 320 can yield lower signal intensity (appearing darker in images) due to a reduction in the detected transverse magnetization Mxy.
• the MR signal from an image can be therefore dependent on the combination of the intrinsic tissue properties and extrinsic user-selected imaging parameters and contrast agents.
• the data processing system 105 passes the selected pulse sequence information to the RF control unit 135 and the gradient control unit 125, which collectively generate the associated waveforms and timings for providing a sequence of pulses to perform a scan.
• the sequence of RF pulses and gradient waveforms, namely the type of pulse sequence, applied may change which relaxation times have the most influence on the image characteristics.
• T2 * relaxation has a significant influence following a 90°RF pulse which is used in a gradient-echo (GRE) sequence
• T2 relaxation has a more significant influence following 90°-180°sequential RF pulses (also known as a spin echo sequence).
• an illustrative pulse sequence 400 is shown that can be used to acquire images using the DREMR system 00. Specifically, a timing diagram for the example pulse sequence is indicated. The timing diagram shows pulse or signal magnitudes, as a function of time, for transmitted (RFt) signal, magnetic field gradients G x , G y , and G z , received RFx signal and filed-shifting signal (FS).
• RFt transmitted
• G x magnetic field gradients
• G y magnetic field gradients
• G z received RFx signal
• FS filed-shifting signal
• An idealized pulse sequence can contain a slice selection radio frequency pulse 410 at RFt, a slice selection gradient pulse 420 at Gz, a phase encoding gradient pulse 430 at Gy, a frequency encoding gradient pulse 440 at Gx, as well as a detected MR signal 450 at RFx.
• the pulses for the three gradients Gx, Gy, and Gz represent the magnitude and duration of the magnetic field gradients that can be generated by the gradient coils 120.
• the slice selection pulse 410 can be generated by the transmit aspect of RF coils 130.
• Detected MR signal 450 can be detected by the receive aspect of the RF coils 130.
• transmit aspect and receive aspect of RF coils 130 are formed by distinct coils.
• the field-shifting signal FS causes the main magnetic field strength to be changed for the duration of the signal FS.
• the precise timing, amplitude, shape and duration of the pulses or signals may vary for different imaging techniques.
• field-shifting signal FS may be applied at a time and manner that allows image contrast to increase for the technique used.
• the first event to occur in pulse sequence 400 can be to turn on the slice selection gradient pulse 420.
• the slice selection RF pulse 410 can be applied at the same time.
• the slice selection RF pulse 410 can be a sine function shaped burst of RF energy. In other implementations, other RF pulse shapes and durations can be used.
• the slice selection gradient pulse 420 can also be turned off and a phase encoding gradient pulse 430 can be turned on.
• the field-shifting signal 460 may also be turned on at this point to change the main magnetic field strength.
• phase encoding gradient pulse 430 is turned off, a frequency encoding gradient pulse 440 can be turned on and a detected MR signal 450 can be recorded.
• shape, magnitudes and durations of the pulses and signals shown in FIG. 4 are chosen for illustrative purposes, and that in implementations, one or more of these factors and others may be varied to achieve the desired scan results.
• the pulse sequence 400 can be repeated a certain number of times or iterations, typically 256 times, to collect all the data needed to produce one image.
• the time between each repetition of the pulse sequence 400 can be referred to as the repetition time (TR).
• the duration between the center point of the slice selection pulse 410 and the peak of detected MR signal 450 can be referred to as echo time (TE). Both TR and TE can be varied as appropriate for a desired scan.
• FIG. 2 is referred to in conjunction with FIG. 4.
• the slice selection gradient pulse 420 can be applied along the Z-axis, satisfying the resonance condition for the protons located in the slice 205.
• the location of the slice along the Z-axis can be determined based in part on the slice selective gradient pulse 420. Accordingly, the slice selection pulse 410, generated at the same time as the slice selection gradient pulse 420 can excite protons that are located within the slice 205 in this example.
• Protons located above and below the slice 205 are typically not affected by the slice selection pulse 410.
• a phase encoding gradient pulse 430 can be applied after the slice selection gradient pulse 420. Assuming this is applied along the Y-axis, the spins at different locations along the Y-axis can begin to precess at different Larmor frequencies. When the phase encoding gradient pulse 420 is turned off, the net magnetization vectors at different locations can precess at the same rate, but possess different phases. The phases can be determined by the duration and magnitude of the phase encoding gradient pulse 430.
• a frequency encoding gradient pulse 440 can be turned on.
• the frequency encoding gradient is in the X direction.
• the frequency encoding gradient can cause protons in the selected slice to precess at rates dependent on their X location. Accordingly, different spatial locations within the slice are now characterized by unique phase angles and precessional frequencies.
• RF receive coils 30 can be used to receive the detected signal 450 generated by the protons contained in the object being scanned while the frequency encoding gradient pulse 440 is turned on.
• the acquired signals can be stored in a temporary matrix referred to as k-space, as shown in FIG 5 at 500.
• k-space is the collection of the detected signals measured for a scan and is in the spatial frequency domain.
• K-space can be covered by frequency encoding data along the X-axis 520 (Kx) and phase encoding data along the Y-axis 530 (Ky).
• Kx frequency encoding data along the X-axis 520
• Ky phase encoding data along the Y-axis 530
• k-space can hold raw data before reconstruction of the image into the spatial domain.
• k-space has the same number of rows and columns as the final image and is filled with raw data during the scan, usually one line per pulse sequence 400.
• the first line of k-space 500 is filled after the completion of the first iteration of the pulse sequence generated for scanning a slice and contains the detected signal for that pulse sequence iteration.
• the k-space can be filled.
• Each iteration of the pulse sequence may be varied slightly, so that signals for the appropriate portions of the k-space are acquired. It should be noted that based on different pulse sequences, other methods of filling the k-space are possible, such as in a spiral manner, and are contemplated.
• the gradient coils 120 produce time-varying magnetic fields with a specific spatial distribution and are a typical component of MRI systems. Greater field-variation magnitudes enable faster MR imaging sequences and increased resolution. As discussed above, the maximum achievable gradient strength is characterized by the gradient efficiency.
• the efficiency of the gradient coils 120 can be improved through variations in the shape, size and placement of the gradient coils 120. For example, in a cylindrical implementation the primary gradient coil windings may be built at a smaller radius closer to the object in the imaging volume. Alternatively, the number of wires (winding density) can be increased.
• the field- shifting electromagnets 140 produce time-varying magnetic fields that can augment the main magnetic field produced by the main magnet 1 0. Greater magnitude of the field variation can enable increased performance. As discussed above, the maximum achievable field-shifting amplitude is characterized by the field-shifting efficiency. The efficiency of the field-shifting electromagnets 140 can be improved through variations in the shape, size and placement of the field-shifting electromagnets 140. For example, the primary field-shifting coils may be built at a smaller radius closer to the object placed in the imaging volume. Alternatively, the number of wires (winding density) can be increased.
• Combining field-shifting electromagnets 140 in the same mechanical assembly as the gradient coils 120 may enable increasing field-shifting efficiency within a given radial space.
• the primary and the shield coils of the gradient- coils 120 may be combined with the primary and shield coils of the field-shifting electromagnets 140 to form a single integrated magnet device with a layer placement that optimizes field-shifting efficiency.
• FIG. 6 Cross section of an example cylindrical implementation for the integrated magnet device 600 is indicated in FIG. 6, in a simplified manner for illustrative purposes.
• the elements shown in FIG. 6 are not to scale.
• the integrated magnet device of this example has a cylindrical shape whose length runs in a plane perpendicular to the plane of the figure. In other implementations, other shapes are possible, as long as the ordering of the layers is preserved. In further variations, the ordering may change as well.
• the integrated magnet device 600 of this example may surround an imaging volume 650, and may include primary field-shifting magnets 605, primary gradient magnets 610, gradient shield magnets 615 and field-shifting shield magnets 620.
• the integrated device is shown within the bore of the magnet 110, indicated at 625.
• RF coils 130 and other magnets or coils such as shim coils may also be placed within imaging volume 650.
• typically primary field-shifting magnets 605 and field-shifting shielding magnets 620 can be used to produce a field- shifting magnetic field along the Z-axis, which can augment the main magnetic field by a predetermined amount dB.
• the Z-axis is perpendicular to the plane of the figure.
• primary field-shifting magnets 605 and field-shifting shielding magnets 620 can include appropriate windings to produce a field-shifting magnetic field along the Z-axis.
• Gradient primary magnets 610 and gradient-shielding magnets 615 can produce fields that vary along all three orthogonal axis X, Y (located in the plane of the figure) and Z, and thus can include windings appropriate for generating and shielding gradient fields along these directions.
• the primary magnets 605 of the field-shifting electromagnets 140 can be placed as close as possible to the object being scanned while the field-shifting shield magnets 620 can be placed farther away from the primary magnets 605 of the field-shifting electromagnet 140, increasing the field-shifting efficiency.
• field-shifting efficiency is prioritized over gradient efficiency by placing the gradient coils 120 in between the primary field-shifting magnets 605 and field-shifting shield magnets 620 of the field-shifting electromagnets 140.
• Additional field-shifting electromagnet 140 layers may be inserted between the innermost primary electromagnet layer 605 and outermost shield electromagnet layer 620 of the field-shifting electromagnet, for example in order to increase the efficiency of the field shifting electromagnet 140 or as part of the field-shifting shield magnet. It should be noted that this layer placement is applicable to differently sized DREMR systems 100 such as DREMR systems 100 scaled for both small animal and human use.
• Thermal power dissipation for gradient coils 120 and field-shifting magnets 140 can be managed by active and passive cooling. Heat can be extracted directly by using conductors having a hollow channel through which coolant is passed, or indirectly by passing coolant through the magnet or coil assembly in a manner such that the coolant is in thermal contact with the windings, or in any other way that is capable of extracting heat energy at the same average rate that resistive power is being dissipated by the electromagnets. Efficiencies can be gained when a cooling layer is in thermal contact with multiple coil components. The same cooling layers can be used to cool the gradient coils and their shields as well as the field-shifting primary and shield coils, making better use of the available radial space.
• FIG. 7 a simplified illustrative cross section of an example integrated magnet device 600 in accordance with the layer placement indicated at FIG. 6 along A-A.
• the example integrated magnet device includes the shielded field-shifting electromagnets 140, the shielded gradient coils 120 and passive and active cooling.
• 650 indicates the imaging volume into which the object to be scanned would be placed.
• RF coils 130 and other coils such as shim coils may also be placed within this space.
• 625 indicates the bore of the main magnet 1 10 within which the integrated magnet device is located.
• substrate layers are indicated at 705.
• the substrate layers can be formed of any rigid or semi-rigid material which can provide mechanical support for the field-shifting electromagnets 140.
• the substrate 705 can be formed of G10, FR4 or epoxy.
• Primary field-shifting magnets of the field-shifting electromagnet 140 are indicated at 710.
• the primary field-shifting magnets 7 0 can be placed on and/or in the substrate 705 and form a magnet that produces the field-shifting magnetic field when activated.
• the primary field-shifting magnets 710 can be formed of windings made from electrically conductive materials suitable for magnetic field generation such as copper.
• the electrically conductive materials used are typically insulated to prevent short circuits within the windings as well as with other nearby electrically conductive components.
• Thermally conductive sub-layers 715a and 715c as well as a coolant sub-layer 715b form a first cooling layer 715.
• the thermally conductive sub-layers 715a and 715c can be formed of any thermally conductive materials such as copper, aluminum, steel (typically uninsulated) or thermally conductive epoxy.
• the coolant sub-layer 7 5b may be composed of any mechanisms which would allow a coolant in the form of a liquid such as water or a glycol mixture or gas such as air to be circulated about the primary field-shifting windings 710 and the primary windings of the gradient coils 120, which are discussed further below.
• hundreds of thin coolant tubes running the length of the primary field-shifting magnets 710 may be used to circulate the liquid coolant.
• the coolant tubes may run around the circumference of the primary field-shifting magnets 710.
• the coolant mechanism used to distribute a coolant in the coolant layer 715b is not electrically conductive.
• the coolant tubes may be constructed from non-conductive material, or may be rendered non-conductive through application of electrically insulating materials. It should be noted that radial space may be conserved by using each cooling layer to cool several coil layers.
• the next three layers of the example integrated magnet device, 725a, 725b and 725c include the primary magnets for gradient coils 120 for producing gradients along the X, Z and Y directions respectively.
• the order of placement of the orthogonal portions of the gradient coils 120 is not limiting and can be varied.
• the primary gradient magnets for producing a gradient along the Z-axis can be placed at 725a
• the magnets for producing a gradient along the Y-axis can be placed at 725b
• the magnets for producing a gradient along the X-axis can be placed at 725c.
• the magnets can be formed from windings that can be made of electrically conductive materials suitable for magnetic field generation such as copper.
• the electrically conductive materials used are typically insulated to prevent short circuits within the windings as well as with other nearby electrically conductive components.
• the primary magnets of the gradient coils 120 are typically placed in and/or on thermally conductive substrates such as epoxy.
• one or more of the primary magnets may be placed in and/or on a rigid or semi-rigid substrate to increase mechanical stability of the layers.
• potting epoxy may be used to form a mechanically stable structure filling all voids between and around the primary windings of the gradient coils 120.
• the epoxy substrate should be thermally conductive in order to efficiently transfer the heat from the primary gradient windings to the coolant layer.
• at least one of the substrates for the three layers may not be thermally conductive, and be formed of materials such as G10 and FR4.
• layer 725b is formed from a thermally non-conductive substrate.
• thermally conductive sub-layers 730a and 730c as well as a coolant sub-layer 730b form a second cooling layer 730.
• the thermally conductive sub-layers 730a and 730c can be formed of any thermally conductive materials such as copper, aluminum, steel or epoxy.
• the coolant sub-layer 730b may be composed of any mechanisms which would allow a coolant in the form of a liquid such as water or gas such as air to be circulated about the primary windings of the gradient coils 120 and the return layer 740, which is discussed further below.
• hundreds of thin coolant tubes running the length of the primary gradient coils 120 may be used to circulate the liquid coolant.
• the coolant tubes may run around the circumference of the primary gradient coils 120 and/or its component magnets.
• the coolant mechanism used to distribute a coolant in the coolant sub-layer 730b is not electrically conductive.
• the coolant tubes may be constructed from non-conductive material, or may be rendered non-conductive through application of insulating materials.
• the direction of coolant flow may vary from that of the coolant layer 7 5. It should be noted that radial space may be conserved by using each cooling layer to cool several magnet or coil layers.
• a layer for return wires and/or return cooling lines is indicated.
• the return wires allow the output current from the field-shifting coils 140 and gradient coils 20 to their respective power supplies to be on the same side of the coil system as the input current. These wires are not actively involved in the production of magnetic fields used for scanning.
• the wires are typically embedded in and/or on a thermally conductive substrate such as potting epoxy.
• Return cooling lines allow the outlet of the cooling system to be on the same side of the coil system as the inlet of the cooling system.
• the return wires and/or the return cooling lines may be arranged in such a way that they do not pass through the space provided by layer 740.
• the next three layers of the example integrated magnet device, 750a, 750b and 750c include the shield magnets for gradient coils 120 for producing shields along the X, Z and Y axis respectively.
• the order of placement of the directional portions of the gradient coils 120 is not limiting and can be varied.
• the shield windings for producing a shield along the Z-axis can be placed at 750a
• the windings for producing a shield along the Y-axis can be placed at 750b
• the windings for producing a shield along the X-axis can be placed at 750c.
• the order can be matched to the order of the primary magnets of the gradient coils 120 in layers 725a through 725c.
• the shield magnets for gradient coils 120 can be formed of windings that can be made from electrically conductive materials suitable for magnetic field
• the electrically conductive materials used are typically insulated to prevent short circuits within the windings as well as with other nearby electrically conductive components.
• the shield magnets of the gradient coils 120 are typically placed in and/or on thermally conductive substrates such as epoxy. In variations, one or more of the primary windings may be placed in and/or on a rigid or semi-rigid substrate to increase mechanical stability of the layers. In some
• At least one of the substrates for the three layers may not be thermally conductive, and be formed of materials such as G10, FR4 or epoxy.
• layer 750b is formed from a thermally non-conductive substrate.
• thermally conductive sub-layers 760a and 760c as well as a coolant sub-layer 760b form the third cooling layer 760.
• the thermally conductive sub-layers 760a and 760c can be formed of any thermally conductive materials such as copper, aluminum, steel or epoxy.
• the coolant sub-layer 760b may be composed of any mechanisms which would allow a coolant in the form of a liquid such as water or a glycol mixture or gas such as air to be circulated about the shield magnets of the gradient coils 120 and the shield magnets of the field-shifting electromagnet 140 layer 770.
• the coolant tubes may run around the circumference of the primary gradient coils 120.
• the coolant mechanism used to distribute a coolant in the coolant sub-layer 760b is not electrically conductive.
• the coolant tubes may be constructed from non-conductive material, or may be rendered non-conductive through application of insulating materials.
• the direction of coolant flow may vary from that of coolant layer 715 and/or 730. It should be noted that radial space may be conserved by using each cooling layer to cool several coil layers.
• Shield magnets of the field shifting electromagnet 140 are indicated at 770.
• the shield magnet layer 760 can be placed on and/or in a substrate such as G10, FR4 or epoxy and form an electromagnet that produces the shield for the field shifting magnet field when activated.
• the field-shifting shield windings can be made from electrically conductive materials suitable for magnetic field generation such as copper.
• the electrically conductive materials used are typically insulated to prevent short circuits within the windings as well as with other nearby electrically conductive components.
• one or more layers of the integrated magnet device may be omitted and/or varied and/or additional layers may be added.
• the three cooling layers are indicated as being formed from substantially the same mechanism for cooling having substantially the same components, in variations, one or more of the three cooling layers 715, 730 and/or 760 can use different cooling mechanisms, sub-layers and/or components from each other.
• other methods of cooling more suitable for a larger DREMR system 100 such as one for human scale applications, can be used.
• hollow coolant- carrying conductors can be used to implement the field-shifting electromagnets 140 and/or gradient coils 120, including the shield windings, and coolant fluids can be circulated through the hollow conductors.
• additional layers of field-shifting electromagnet conductors, gradient coil conductors or other electromagnet conductors may be inserted within the assembly.
• the substrates in which the magnets are placed may be hold-offs spaced around the circumference of the magnet to hold the wire for some of the electromagnets.
• the hold- offs may be printed using a three dimensional printer.
• the integrated magnet device 600 can be used for DREMR systems 100 of a wide range of geometries and sizes.
• DREMR system 00 can be constructed using the integrated magnet device in appropriate sizes and shapes to accommodate human-scale imaging applications, such as brain imaging applications for scanning the brain, or for small animal scanning applications.
• the relative radial positions or ordering of the windings in the integrated magnet device 600, including the shield magnets for field-shifting electromagnets 140 and the gradient coils 120, can remain the same regardless of the size of the DREMR system 100.
• the cooling methods and mechanisms used may be varied.
• hollow conductors for the electromagnet windings can be used to provide coolant flow that is placed closer to the heat source. Accordingly, coolant fluid, whether liquid or gas, would flow through the hollow conductors of the windings forming the magnets, in place of or in addition to the coolant layers of the integrated magnet device.
• hollow conductors can run in a loop, a spiral or helix around a radius of the integrated magnet device to implement the longitudinal (z-axis) gradient electromagnet of the gradient coils 120 and a transverse (x-axis or y-axis) gradient wire pattern can be placed in thermal contact with the z layer that contains the coolant flow.
• the same method can be utilized to cool the windings of field-shifting electromagnets 1 0, which can be similar to the z-gradient windings in that they run in loops, a spiral or a helix around a radius of the integrated magnet device.
• the transverse (x-axis or y-axis) gradient electromagnet could also be formed of hollow conductors through which coolant fluid, whether liquid or gas, would flow.
• the patterns used to implement windings of the field-shifting electromagnets 140 and the gradient coils 120 may vary. For example, different winding patterns may be used to implement the field-shifting electromagnets 140 and the gradient coils 120 to implement an integrated magnet device for use in a human-scale DREMR system 100 intended for scanning brains, in comparison to that which is intended for use in scanning small animals.
• the relative radial ordering of the windings used to implement the field- shifting electromagnets 140 and gradient coils 120, including the shield windings, in the integrated magnet device may remain the same.
• the winding patterns for the field-shifting electromagnets 140 and the gradient coils 120 may not be symmetric along the longitudinal z-axis because a head might not be able to be placed in the center of the magnet due to geometric constraints.
• the winding patterns for the field-shifting electromagnets 140 and the gradient coils 120 may not be symmetric along the longitudinal z-axis because a head might not be able to be placed in the center of the magnet due to geometric constraints.
• electromagnets 140 and the gradient coils 120 for a DREMR system 00 intended for small animal imaging may be longitudinally symmetric because the imaging region can be in the center of the coil.

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